A major limitation to the clinical goal of achieving ideal diabetic glucose control is the unavailability of unlimited and/or continuous glucose monitoring. Despite the non-invasive advances described in U.S. Pat. No. 4,975,581 to Robinson, et al., a lancet cut into the finger is still necessary for all present forms of home glucose monitoring. This is so compromising to the diabetic patient that the most effective use of any form of diabetic management is rarely achieved, including multiple insulin shots, continuous subcutaneous pump delivery, intraperitoneal or intravascular implanted pump delivery, or oral diabetic pharmaceutical agents. It is possible that diabetic glycemia could be controlled with conventional treatment, external pumps, or implanted insulin delivery devices, if on-line or continuous glucose levels were known by the patient or by a monitoring system. Such information would enable development of a closed loop insulin delivery system.
The theoretical basis for non-invasive glucose determination is based upon quantitative infrared spectroscopy. Infrared spectroscopy measures the electromagnetic radiation (0.7-25 .mu.m) a substance absorbs at various wavelengths. Molecules do not maintain fixed positions with respect to each other but vibrate back and forth about an average distance. Absorption of light at the appropriate energy causes the molecule to become excited to a higher vibrational level. The excitation of the molecule to an excited state occurs only at certain discrete energy levels, which are characteristic for that particular molecule. Most primary vibrational states occur in the mid-infrared frequency region (i.e., 2.5-25 .mu.m). However, noninvasive analyte determination in this region is problematic, if not impossible, due to the absorption of the light by water. The problem is overcome through the use of shorter wavelengths of light which are not as attenuated by water. Overtones of the primary vibrational states exist at shorter wavelengths and enable quantitative determination at these wavelengths. Overtones of the primary vibrations occur at 1/2, 1/3, 1/4 . . . and so on of the wavelength of the fundamental mode. Additionally, combination bands also exist. A combination band occurs when the radiation has the correct energy to excite two vibrations at once.
Although glucose absorbs at multiple frequencies in both the mid and near infrared, there are other infrared active analytes in the blood which also absorb at similar frequencies. Due to the overlapping nature of these absorption bands no single or specific frequency can be used for reliable noninvasive glucose measurement. Analysis of spectral data for glucose measurement thus requires evaluation of many spectral intensities over a wide spectral range to achieve the sensitivity, precision, accuracy, and reliability necessary for quantitative determination. This is also true for other blood analytes. In addition to overlapping absorption bands, measurement of glucose is further complicated by the fact that glucose is a minor component by weight in blood and that the resulting spectral data may exhibit a nonlinear response due to both the properties of the substance being examined and/or inherent nonlinearities in optical instrumentation.
The difficulty of modeling the spectral response requires, as set forth in U.S. Pat. No. 4,975,581, the use of multivariate statistical methods rather than univariate methods. These techniques allow information to be extracted from data which cannot be obtained by other data analysis routines. The methods previously disclosed in U.S. Pat. No. 4,975,581, increase analytical precision to the point where the spectroscopic methods become useful for clinical determinations.
Using expensive optical instrumentation, the technology disclosed in U.S. Pat. No. 4,975,581 has been applied for the quantitative measurement of analytes in biological fluids. The focus of this effort has been in the area of noninvasive glucose measurement, portions of which are described in: (1) "Post-Prandial Blood Glucose Determination by Quantitative Mid-Infrared Spectroscopy", K. J. Ward, D. M. Haaland, M. R. Robinson and R. P. Eaton, Applied Spectroscopy, Vol. 46, No. 6, 1992, pages 959-965, (2) "Reagentless Near-Infrared Determination of Glucose In Whole Blood Using Multivariate Calibration", D. M. Haaland, M. R. Robinson, G. W. Koepp, E. V. Thomas, and R. P. Eaton, Applied Spectroscopy, Vol. 46, No. 10, 1992, pages 1575-1578, and (3) "Noninvasive Glucose Monitoring in Diabetic Patients: a Preliminary Evaluation", M. R. Robinson, R. P. Eaton, D. M. Haaland, G. W. Koepp, E. V. Thomas, B. R. Stallard and P. L. Robinson, Clinical Chemistry, Vol. 38, No. 9, 1992, pages 1618-1622.
In addition to the body of glucose research disclosed by the foregoing papers, M. K. Alam, R. P. Eaton, D. M. Haaland, M. R. Robinson, P. L. Robinson and E. V. Thomas (hereinafter Alam, et. al.) have worked extensively in the near infrared from 700 to 1400 nm. This spectral region allows transmission of the infrared light through the finger and contains meaningful glucose information. With type-I diabetic volunteers, three representative instrument configurations were investigated. In the first, as disclosed in Clinical Chemistry, Vol. 38, No. 9, a Nicolet 800 FTIR instrument equipped with a InSb detector was used. The second system utilized a SPEX grating spectrometer equipped with a germanium array detector. In the third configuration, the SPEX grating spectrometer equipped with the germanium array detector was coupled with fiber optics, which transmitted the light from the instrument to the finger and from the finger back to the instrument. The clinical protocol and method for evaluation of IR spectroscopy for the in vitro determination of blood glucose is described in more detail in the above identified patent and published papers. Work on the second and third configurations has not been published.
With a Nicolet 800 Fourier transform infrared spectrometer (FTIR) equipped with an InSb detector, a diabetic patient undergoing a meal tolerance test was examined using near-infrared transmission measurements through his finger. The patient's blood glucose levels varied between 48 mg/dl and 481 mg/dl, with 41 samples obtained. The average absolute error of prediction on all samples was 19.8 mg/dl. The data are plotted in FIG. 1.
The feasibility of non-invasive glucose determination was next investigated on a grating spectrometer equipped with a germanium array detector. The optical sampling method was transmission of light (800-1330 nm) through the patient's index finger. The patient's blood glucose level varied between 92 mg/dl and 434 mg/dl, with 29 samples obtained. The average absolute error of prediction for this data was approximately 24.3 mg/dl. The data are plotted in FIG. 2.
In the final instrument configuration, the grating spectrometer-germanium detector instrument was outfitted with a fiber optic sampling configuration. Fiber optics were used both to transmit light to the finger and to collect light from the opposite side of the finger. The patient's blood glucose level varied from 83 mg/dl and 399 mg/dl, with 21 samples obtained. Analysis of the data yielded an average absolute error of 11.9 mg/dl. The data are plotted in FIG. 3. The accuracy of this non-invasive determination is comparable to the accuracy of existing invasive home glucose monitors. The results from the fiber optic study were, vis-a-vis the first two configurations, improved due to the ability to repeatedly position the finger between the fiber bundles. Repeatable positioning of the finger decreased the baseline variation observed in the spectra and, it is believed, improved the accuracy of the noninvasive prediction.
In all the above studies, the sampling apparatus used consisted of a circular tube which matched the approximate size of the patients' fingers. The light entered the finger on the palmar side and exited through the fingernail. Although clinically useful measurements were made, the finger sampling techniques used and the instrumentation employed during these studies are not optimal, extremely expensive, and not suitable for either clinical or home use. Thus, improvements in both areas are required before a device can be made available for use by the diabetic patient.
In general, the sampling device should perform two major functions:
1. Enable maximal procurement of spectral information for measurement of the analyte of interest; and PA1 2. Minimize those spectral variations associated with sampling the tissue that adversely influence the quantitative measurement of the analyte. PA1 1. The sampling apparatus utilized does not allow measurement of any wavelengths containing glucose information in the 1400 to 2400 nm region; PA1 2. The sampling apparatus does not optimize sampling geometry for the light propagation characteristics of the wavelengths to be measured; PA1 3. There is no compensation for the influence of skin pigmentation differences between patients; PA1 4. The finger sampling apparatus utilized does not allow repeated sampling of a single patient's finger and does not minimize between patient differences; PA1 5. There is no compensation for the influence of arterial pulsations in patients' tissue (e.g., finger); and PA1 6. The sampling device is not temperature controlled. PA1 1. Which enables maximal use of the various wavelength regions that include spectral information on the blood analyte of interest, (e.g. glucose); PA1 2. Which optimizes the path, depending upon the propagation characteristics of the wavelength used; PA1 3. Decreases or compensates for the influence of those substances present in the body that exhibit spectral overlap and have varying concentrations; PA1 4. Which reduces the spectral variability between people and allows for more accurate analyte measurement; PA1 5. Decreases the effects of arterial pulsations; PA1 6. Reduces the affect of body part temperature differences through a thermostated sampling device; PA1 7. Which is rugged and does not require frequent recalibration; and PA1 8. Which uses acousto-optic tunable filters (AOTFs), or other suitably rugged optical instrumentation.
The following specific inadequacies have been identified in prior art sampling devices:
To understand the inadequacies of the current sampling device and associated instrumentation, and to recognize the benefits of the disclosed invention, a general understanding of infrared spectroscopy and of light propagation characteristics in tissue is necessary.
Spectroscopic information which facilitates the measurement of glucose occurs over the majority of the near infrared region. In those wavelength regions suitable for noninvasive measurement glucose has absorption peaks in the following areas: 950-1050 nm; 1150-1300 nm; 1510-1850 nm; and 2070-2370 nm. If correctly processed, the judicious use of spectroscopic information from the entire wavelength region will yield better quantitative results than only one wavelength region. The utility of using all possible information is especially true in complex environments such as human tissue.
To demonstrate the utility of using multiple wavelength regions a simple experiment was performed. A set of cuvette samples containing water, urea, and glucose were optically sampled over the entire wavelength region from 700 to 2400 nm. The optical pathlength used for data acquisition was 1 cm in the 700 to 1400 nm range and 1 mm in the 1400 to 2400 nm range. The change in pathlength was necessitated due to differences in the absorbance of water in given wavelength regions, the importance of which is discussed at length in the Description of the Preferred Embodiments. The resulting spectra was processed using four different wavelength ranges. The results are shown in Table I below.
TABLE I # of # of # of # of Total wavelengths wavelengths wavelengths wavelengths number of Wavelength used in used in used in used in wavelengths Standard error Region 700-1100 nm 1100-1400 nm 1400-2000 nm 2000-2400 nm used of prediction 700-1100 nm 7 N/A N/A N/A 7 18.1 700-1400 nm 6 2 N/A N/A 8 12.2 700-2000 nm 4 4 6 N/A 14 10.4 700-2400 nm 3 7 8 5 23 5.83
Thus, it is clear that the inclusion of information from all wavelength regions containing glucose information improves the accuracy of the optical measurement. Therefore, it is an object of the present invention to provide for a tissue sampling device and associated instrumentation which enables use of spectral information from, in the case of glucose, the entire wavelength region from 700 to 2400 nm.
In order to access the entire wavelength region from 700 to 2400 nm, the light propagation characteristics in tissue at these wavelengths must be understood. Although multiple papers have been published on the optical properties of skin, literature on the optical properties of the entire finger is sparse. The finger is a complex, dynamic, variable, heterogenous and multilayered optical media, which makes entirely rigorous analysis difficult. As light enters the finger it undergoes multiple and diverse scattering effects throughout its tortuous path of propagation. The overall optical response is a combination involving molecular (Rayleigh) scattering, particle (Mie) scattering, and index (Fresnel and Christiansen effect) scattering. A simplistic diagram of the light propagation characteristics of light within the body is schematically illustrated in FIG. 1 of "The Optics of Human Skin", R. Rox Anderson, B. S. and John A. Parrish, M. D., The Journal of Investigative Dermatology, 77:13-19, 1981. For convenience, a slightly modified version is illustrated in FIG. 4. The light entering the tissue is either absorbed, reflected or transmitted. Transmission is defined as that fraction of the radiation incident on one side of the sample that passes through the sample. A second type of optical sampling, "partial transmission" (sometimes referred to as "diffuse reflectance"), is defined as that fraction of light that interacts with the tissue and where the sampling does not require location of the source and detector on opposite sides of the body part. See FIG. 4. "Simple reflectance" is defined as that fraction of radiation incident upon one side that returns directly from the surface of the sample. The information content of this reflected light is negligible as the light is reflected by the bloodless epidermis.
The optical characteristics of the epidermis and dermis are well characterized. Over the region from 400 to 1300 nm, the skin can be modeled by considering the thin epidermis to be an optically absorbing element with negligible scattering, overlying the thick dermis, which acts as a diffuse reflector. Transmission through the epidermis is mainly a function of melanin, which resides solely in the epidermis. The transmission characteristics of the dermal element depends upon both scattering mainly by collagen and absorption. It is important to note that in both transmission and partial transmission the light must transverse the melanin containing epidermis twice. Returning to FIG. 4, it will be seen that light penetrates the epidermis, interacts with the dermis, and will eventually interact with subcutaneous tissues. The capillary bed below the epidermis is the most superficial vascular layer and light propagation to this level or deeper is desired for reliable non-invasive analyte determination. Thus, light having transversed this region will contain the necessary spectral information for analyte measurement (e.g., alcohol, cholesterol, BUN (blood urea nitrogen), creatine, hemoglobin and bilirubin).
As previously demonstrated the use of information from all wavelength regions improves both the sensitivity and specificity of the optical measurement. However, transmission measurements through the finger become problematic at wavelengths greater than 1400 nm due to the absorbance of the radiation by water. Water peaks are seen at 760, 1000, 1200, 1450, and 2000 nm, with each associated band exhibiting a marked increase in absorption. See "Near-infrared Studies of the Structure of Water. I. Pure Water", Buijs, K. and Choppin, G. R., Journal of Chemical Physics, Vol. 39, No. 8, October 1965. The human body is approximately 70% water and in the near infrared spectral region water is the largest absorber. When considering the finger, it can be simplistically modeled as a highly scattering aqueous media surrounded by skin. FIG. 5 shows a simplified model of a finger/thumb 11, which can be grossly modelled as a water cuvette having a pathlength of 1 cm. Finger/thumb 11 includes epidermis 13, dermis 15, the subcutaneous tissue 17 and bone 19. FIG. 6 shows the absorbance of water versus wavelength for a 1 mm pathlength on the right hand axis and shows relative pathlength versus wavelength on the left hand axis. Although influenced by the intensity of the light source, the relative pathlength on the y-axis corresponds to a unit absorbance of one at the x-axis wavelength location. Examination of FIG. 6 reveals that transmission measurements through the finger are difficult at wavelengths of greater than 1400 nm using a standard tungsten halogen lamp due to limitations in light propagation. These concepts concerning optical transmission characteristics in water are also shown in FIG. 11 of U.S. Pat. No. 4,570,638. Note, the intensity of the light source will influence the relative pathlengths at different wavelengths. The pathlengths referenced above are appropriate for a standard 100 watt tungsten halogen source.
Despite the absorbance by water, partial transmission sampling can be performed in a tissue at wavelengths greater than 1400 nm. For example, the light may enter the tissue and exit the tissue several millimeters away, as shown in FIG. 4. Thus, the light has been transmitted but only through a portion of the tissue or body part. It is an object of the present invention to use partial transmission for the procurement of spectral information at wavelengths longer than those which lend themselves to standard transmission sampling.
In quantitative spectroscopy it is desirable to maximize pathlength through the sample while maintaining an adequate signal-to-noise ratio at the detector. Thus, when making measurements in the 1000 nm region, spectral data obtained from a 1 cm cuvette will typically outperform data obtained from a 1 mm cuvette if the two sets of data have similar signal-to-noise ratios. The improved performance occurs due to the longer pathlength which forces the light to-interact with the absorbing substance for a longer period of time. Thus, it is an object of the present invention to optimize the path for given wavelength regions.
The procurement of maximal spectral information can be augmented by sampling through the nail of the finger. In simplistic terms the nail provides a "window" to the highly vascular nail bed, similar to scraping away the upper layers of a person's skin and placing a glass slide on the resulting surface. This "window" is the result of the fact that optical penetration through the nail is greater than the skin, Physical Properties of Tissue, Chapter 3, Francis A. Duck, Academic Press 1991. Transmission differences between the nail and tissue become greater at increasing wavelength due to hydration differences between the nail and skin. As the nail has a lower water content than the skin, it facilitates spectral sampling in the longer wavelength regions.
The histological structure of the nail and nail bed further facilitates noninvasive sampling. The nail bed is defined as the area of skin covered by the nail. The epithelium of the nail bed is significantly thinner than the normal epidermis. The normal epidermis is composed of five layers: (1) stratum corneum (the outermost layer); (2) stratum lucidum; (3) stratum granulosum; (4) stratum spinosum; and (5) stratum basale (the innermost layer). The nail bed is composed of only the stratum basale and the stratum spinosum, and lacks significant keratinization. Thus, the physical distance to the highly vascular dermis is less in the nail bed. In summary, the nail is highly transmissive, has a low water content, and covers a highly vascular structure which makes optical sampling through the nail region highly desirable. The low water content of the nail is especially desirable when sampling in the 1400 to 2400 nm region. Therefore, it is an object of this invention to utilize this "window" into the body for procurement of maximal analyte information.
As the light passes through a body part such as a finger or thumb, it interacts with both tissue, intracellular fluid, and blood. As the object of the invention is to measure blood analytes, maximizing the amount of blood in the tissue being irradiated should improve the measurement. The accuracy of noninvasive measurement is determined by its correlation to standard invasive blood measurements. As the noninvasive measurement is actually a blood/tissue measurement, use of highly vascular body parts and maximization of blood content will improve measurement accuracy. The fingers and palm have much higher capillary densities than the arms, legs, or trunk, and are thus desired sampling locations.
Blood can be concentrated in the finger by several methods but venous engorgement is a method easily performed. The arterial system operates at a higher pressure than the venous system. Thus, occlusion of the venous system allows the finger to be pumped full of blood by the arterial system. The result is a finger with an above normal amount of blood (i.e., venous engorgement). However, the continued filling of the finger can cause instabilities in the optical measurement as the blood volume of the finger is changing. This change can be minimized by subsequently occluding arterial flow following proper "filling" of the finger. The result is venous engorgement of the finger which, when performed in a repeatable fashion, will enhance noninvasive analyte measurement.
The second major function of the sampling device is to minimize those spectral variations associated with sampling the finger that adversely influence the quantitative measurement of the analyte. The major problems recognized by the Applicants are: skin pigmentation differences between patients, arterial pulsations, finger thickness differences, instabilities in finger sampling, and the lack of temperature control.
Quantitative spectroscopic measurement becomes increasingly difficult as the complexity of the matrix under irridation increases. For example, FIG. 7 shows the absorbance spectra for water, glucose, alcohol and urea in the 900 to 1350 nm region. As can be seen, these substances have different absorbance characteristics, but there exist no single wavelengths where only one analyte absorbs. Spectral overlaps also exist in the wavelength region from 1350-2400 nm. Thus, the ability to isolate a single band for analyte measurement is difficult in an environment of overlapping spectral absorbencies. The degree of spectral overlap associated with noninvasive measurements is significant and any method which diminishes spectral overlap will decrease the complexity of the measurement process.
The problem of overlapping spectral absorbencies is complicated further in the wavelength region from 300 to 1100 nm due to the spectral absorbencies of melanin, bilirubin, and hemoglobin. See FIG. 8, "The Optics of Human Skin," supra. When considering the influence of overlapping spectra, the amount of overlap and the variation in the concentration of the overlapping substance are important parameters. For example, if a given substance only absorbed within a one nanometer bandwidth then omission of that wavelength would resolve the overlap problem completely. However, if the substance has a broad absorbance then omission of the wavelength region in which it absorbs is not reasonable. Thus, compensating for substances with broad spectral absorbencies is especially problematic.
In the wavelength region from 300 to 1100 nm, the optical absorbance of melanin varies significantly due to gross variations in skin pigmentation. Contrary to popular belief, melanin does not absorb light like a "neutral density" filter in the skin. Absorption by melanin decreases steadily from short wavelengths to longer wavelengths and does not have significant absorption above 1100 nm. See "The Optics of Human Skin", supra. Thus, in the wavelength region from 300 to 1100 nm melanin exhibits a broad, varying spectral influence which complicates noninvasive analyte measurement in heterogenous patient populations, especially those with varying ethnicity.
Hemoglobin is also an important absorber in the 300 to 800 nm region because its spectral influence varies as the amount of blood in a given tissue area changes. The modulation of light by pulsatile arterial flow is the fundamental principle upon which pulse oximetry is based. The amount of optical change observed due to arterial pulsations is a function of the patients' overall vascular status, heart rate and pulse pressure. As is the situation with melanin, the varying spectral absorbance of hemoglobin makes it a difficult component to model when developing multivariate calibration models. Thus, it is an object of the present invention to provide a methodology for removing these large spectral changes introduced by arterial pulsations, which do not relate to the concentration of the particular analyte of interest (e.g. glucose).
The spectral variation introduced by arterial pulsations can be minimized by understanding the physiology associated with arterial pulses. The pressure and corresponding pulse size in the arterial system near the heart is quite high. As the blood flows to the periphery and the vessels become progressively smaller both the mean pressure and pulse size decrease. As the blood enters the capillary bed the mean pressure has decreased to less than 40 mm hg and there is no longer any significant pulsatile component to the capillary blood flow. However, the light passing into and out of the tissue interacts with blood in both the pulsing arterioles and the capillary bed. The noise introduced by the pulsing arterioles can be removed by simple compression of the tissue (e.g., finger) or proximal compression of the arterial system. If one presses the finger against a sampling device, or has it compressed externally, the arterial pulsations can be minimized or removed. Thus, the simple removal of the spectral variance resulting from arterial pulsations can improve the signal-to-noise ratio of the resulting spectra, and thus improve the precision of the analyte measurement.
In addition to the foregoing problems associated with pigmentation and arterial pulsations, the quality of any noninvasive spectroscopic measurement will be improved if the sampling conditions are repeatable. In the sampling of human subjects, control of all sampling parameters becomes extremely difficult. The most significant problems identified by Alam et al. are variations in finger thickness and finger temperature.
Variation in finger thickness complicates the process of preforming noninvasive analyte measurements. In the spectroscopic literature, the majority of all quantitative spectroscopy is done with a fixed optical pathlength. In human applications such a requirement becomes impossible to satisfy. The magnitude of the problem can be reduced through the use of partial transmission sampling, often referred to as diffuse reflectance sampling. In partial transmission sampling the mean optical pathlength through the finger is determined in large part by the separation between the source and detector. The separation distance is not the sole influence on mean optical pathlength as differences in tissue composition and other physiological parameters will influence the light propagation. With reference to FIG. 4, the source and detector are on the same side of the tissue during partial transmission sampling. Due to their location on the same side of the tissue, tissue thickness has a reduced influence on the measurement. The mean optical pathlength then becomes a function of the separation between source and detector. As previously stated partial transmission sampling will reduce the spectral variation introduce by differences in tissue thickness.
The spectral variation introduced by differences in tissue temperature also complicates the noninvasive measurement of analytes. J. Lin and Cris W. Brown, "Near-IR Fiber-optic Probe for Electrolytes in Aqueous Solution", Analytical Chemistry, Vol. 65, pages 287-292, 1993, have shown that the near infrared spectral region is sensitive to temperature effects. Marked spectral changes were observed when water solutions were subjected to temperature changes from 20.degree. C. to 35.degree. C. The regions most sensitive to temperature are those having extensive hydrogen bonding. These regions exhibit spectral changes as the hydrogen bonding changes with increasing temperature.
In studies by Alam et al. differences in skin temperature and its influence have been observed. In the prior articles on glucose, the sampling device (constructed of aluminum) was not temperature controlled and therefore acted as a heat sink. The result of using such an unthermostated sampling device is to change the skin temperature of the patient. As mentioned previously, any spectral influence which is broad in nature, not constant, and does not relate to glucose concentration can degrade the accuracy of the glucose measurement. At a minimum, such spectral changes increase the complexity of the multivariate calibration. The problem can be overcome or at least compensated for by thermostating the tissue and the sampling device. In the case of the finger, the average physiological temperature is approximately 82.degree. F. with an average variation of .+-.5.degree. F. The sampling device can be heated to an above normal tissue temperature to increase blood flow to the tissue area in contact with the device. The result is an increase in the vascular supply to the tissue and a corresponding increase in the blood content of the tissue. The end result of temperature regulation is a reduction in spectral variation not associated with glucose and an improvement in measurement accuracy.
In addition to problems directly associated with sampling the finger, the actual optical instrumentation discussed above is not well suited for commercial realization of a noninvasive analyte monitor. For clinical applications the spectrometer must be rugged without the need for frequent maintenance and re-calibration. Fixed grating spectrometers afford multiplex data acquisition and can be suited to the clinical environment such that maintenance and/or re-calibration are minimized. However, the accurate measurement of selected wavelengths is still a function of a precise geometrical arrangement between the grating and the detector. Vibration or mishandling can cause "blurring" of the image on the array, which translates into reduced performance. Fourier Transform spectrometers are available in the near infrared spectral region and are capable excellent resolution and sensitivity. Unfortunately since most FTIR spectrometers require precision translation of mirrors, their performance is also typically sensitive to the environment (i.e., vibrations and dust). Thus, it is highly unlikely that the instrument configurations used to demonstrate feasibility of noninvasive glucose measurement will satisfy the commercial environment.
Spectra can be generated by using multiple band-pass filters. The instruments afford high optical throughput but have limited flexibility, as a separate filter is needed for each wavelength intensity measured. Nevertheless, as disclosed in the Description of the Preferred Embodiments, filter instruments do represent a viable technology suitable for commercial realization.
The above described problems of limited stability and frequent recalibration can be addressed by using an acousto-optic tunable filter (AOTF). AOTFs are solid-state devices which utilize acousto-optic interactions in an anisotropic medium. The result is a compact solid-state spectrometer that can be tuned electronically in a matter of microseconds over a wide spectral range encompassing both the UV and IR regions. Due to its solid-state design, there are no moving parts. The AOTF is therefore immune to orientation changes and even significant shock and vibration. AOTFs are capable of excellent resolution and can be incorporated into sealed systems. The end result is a small, durable light dispersion device which allows random access to different wavelengths. See "Acousto-optic devices", Chieu D. Tran, Analytical Chemistry, Vol 64, No 20 Oct. 15, 1992). Also see; Photonics Global Forecast, Defense-Related Acousto-Optics Transform Commercial Products, R. G. Rosemeier, Photonics Spectra, 83-84, January, 1993; U.S. Pat. No. 4,883,963 to G. J. Kemeny et al.; and U.S. Pat. No. 5,120,961 to K. H. Levin et al.
In the operation of an AOTF, the wavelength of the diffracted light depends upon the frequency of the radio frequency (rf) signal applied to the AOTF. Light with relatively shorter wavelengths will de diffracted from the AOTF when higher rf signals are applied to the filter. For example, 514 nm light is diffracted when a 64 MHz rf signal is applied. Increasing the frequency to 75 MHz changes the diffracted wavelength to 457 nm. Thus, by simply changing the frequency of the rf signal, the operator has random access to any desired wavelength in the UV-IR region.
In addition to the previously stated characteristics, AOTFs have an additional characteristic which make them well suited for noninvasive medical instruments. The first is the ability of the AOTF to modulate the intensity of the diffracted light, (i.e. during operation the diffracted light is the light exiting the AOTF with the proper wavelength characteristics). The power of the applied rf signal can be used to control the intensity of the diffracted light. Thus, AOTFs provide a unique way to maintain the intensity of the light of different wavelengths at a desired level. By incorporating a feedback system into the AOTF driver, the power of the rf signal can be controlled and thus the intensity of light hitting the detector is controlled.
In general the preceding instruments are based on dispersion of a broadband light source with subsequent detection of the separated wavelengths. The quantitative measurement of blood analytes can be performed by the use of a discrete number of wavelengths. Specifically, glucose has been measured in vitro solutions composed of glucose, urea, alcohol and water through the use of 20 discrete groups of contiguous wavelengths. Thus, the use of sources that emit in a narrow wavelength region could be used in combination for analyte measurement. Specifically light emitting diodes (LEDs), laser diodes, or tunable lasers could be used for the noninvasive measurement of blood analytes.
As the time required to make the measurement is an important parameter, any instrument recording wavelengths in a multiplex manner is desired. The recording of more than one wavelength during a given time period will result in a multiplex advantage. Optical multiplexing increases the effective signal-to-noise ratio that can be achieved for detector-noise-limited spectroscopic measurements. For example, consider the case of measuring each wavelength intensity one at a time versus measurement of multiple wavelengths on an array detector. Given the same measurement time the signal-to-noise ratio of the array detector will exceed that of the single wavelength measurement device.
It is important to recognize that any device having either multiple detectors or sources can acquire data in a multiplex manner. When using LEDs or multiple single element detectors, they can be energized using Hadamard transform techniques. Through Hadamard transform optical coding techniques the (theoretical) signal-to-noise ratio gains at constant observing time when compared with conventional point-by-point image scanning, can be as high as 1/2N.sup.1/2, where N is the total number of image resolution elements. The principles of Hadamard transformation are explained in the following articles: "Fourier and Hadamard transform methods in Spectroscopy", by A. G. Marshall, et al., The Journal of Analytical Chemistry Vol. 47, No. 4, pp 491A-504A, April 1975 and "Hadamard Transform Image Scanning", by J. A. Decker, Jr., The Journal of Applied Optics, Vol. 9, No. 6, pp 1392-1395, June. 1970.
With Hadamard transform methods, during operation approximately half of the total number of single wavelength emitting devices are energized for a measurement observation. During the second observation a different set of diodes will be energized. The process continues until N (the number of diodes) observations have been made. The end result is N different observations expressed as N linear equations. The solution of these equations yields the specific intensity value associated with each specific diode. Through the Hadamard approach an improvement of a factor of N/2 in signal-to-noise ratio over the conventional one-diode-at-a-time measurement is achieved because half the diodes are energized during each observation, rather than just one. Thus, in the preferred embodiment the diodes or detectors may be energized via Hadamard transform optical coding techniques to maximize signal-to-noise ratios for a given measurement time.
In view of the foregoing, it is an object of this invention to provide an apparatus and associated methodology for the repeatable procurement of spectral data which can be analyzed for noninvasive measurement of blood analytes. More specifically, the objectives of the current invention are to provide a device: